The present invention relates to gamma cameras generally and more specifically to an apparatus and method for use with a gamma camera for selecting a subset of camera detectors for processing to speed up detector operation and enhance image quality.
Single photon emission computed tomography (SPECT) examinations are carried out by injecting a dilution marker comprising a compound labeled with a radiopharmaceutical into the body of a patient to be examined. A radiopharmaceutical is a substance that emits photons at one or more energy levels. By choosing a compound that will accumulate in an organ to be imaged, compound concentration, and hence radiopharmaceutical concentration, can be substantially limited to an organ of interest. A radiopharmaceutical that emits photons or gamma emissions which are approximately at a single known energy level is chosen. The organ to be imaged will be referred to as an organ of interest and an energy range which approximates the known energy level will be referred to as the marker range.
While moving through a patient's blood stream the marker, including the radiopharmaceutical, becomes concentrated in the organ of interest. By measuring the number of photons emitted from the organ of interest which are within the marker range, organ characteristics, including irregularities, can be identified.
To measure the number of emitted photons, one or more planar gamma cameras are used. After a marker has become concentrated within an organ of interest, a camera is positioned at an imaging angle with respect to the organ of interest such that the organ is positioned within the camera's field of view FOV. The camera is designed to detect photons traveling along preferred paths within the FOV.
A gamma camera consists of a collimator, a scintillation crystal, a plurality of photo multiplier tubes (PMTs) and a camera processor. The collimator typically includes a rectangular lead block having a width dimension and a length dimension which together define the FOV. The collimator block forms tiny holes which pass therethrough defining the preferred photon paths. The preferred paths are unidirectional and perpendicular to the length of the collimator. The collimator blocks emissions toward the crystal along non-preferred paths.
The scintillation crystal is positioned adjacent the collimator on a side opposite the FOV and has an impact surface and an oppositely facing emitter surface. The impact surface defines a two dimensional imaging area A having a length L and a width W. Photons which pass through the collimator impact and are absorbed by the impact surface at impact points. The crystal emitter surface emits light from an emitter point adjacent the impact point each time a photon is absorbed. The amount of light emitted depends on the absorbed photon's energy level.
The PMTs are arranged in a two dimensional array which is positioned adjacent the emitter surface. Light emitted by the crystal is detected by the PMTs which are in the area adjacent the emitter point. Each PMT which detects light generates an analog intensity signal which is proportional to the amount of light detected. When a single photon is absorbed by the crystal, the emitted light is typically absorbed by several different PMTs such that several PMTs generate intensity signals simultaneously. For the purposes of this explanation all intensity signals caused by a single photon will be collectively referred to as a signal set.
The processor receives each signal set and performs a plurality of calculations on each signal set to determine two characteristics of the corresponding photon. First, the processor combines the intensity signals of each signal set to identify the energy level of a corresponding photon. Photons having energies within the marker range will be referred to as events. Only signals corresponding to events are used for imaging. Second, the processor performs a series of calculations in an effort to determine precisely where on the impact surface imaging area A an event occurred. Once impact locations of all events have been identified, the processor uses the impact locations to create an image of the organ of interest which corresponds to the camera imaging angle.
To create a three dimensional image of the organ of interest, a gamma camera can be used to generate a plurality of images from different imaging angles. To this end, the camera is positioned parallel to, and at an imaging angle about, a rotation axis which passes through the organ of interest. The angle is incremented between views so that the plurality of images are generated. The plurality of images are then used to construct pictures of transaxial slices of the torso section using algorithms and iterative methods that are well known to those skilled in the tomographic arts.
With any imaging system there are several different criteria by which to judge system usefulness. Perhaps the two most important criteria for judging system usefulness are imaging speed and the quality of resulting images. For the purposes of this explanation the time required to generate an image at one imaging angle will be referred to as an imaging period and the time required to generate images from several imaging angles to generate a three dimensional image will be referred to as an imaging session.
Imaging speed is important for at least three reasons. First, the likelihood of imaging errors increases as imaging session duration is extended. Ideally imaging should be performed while a patient remains completely still. Patient movement can result in blurred images which are unusable for diagnostic purposes. Patient movement is more likely during extended sessions than it is during abbreviated sessions.
Second, speedy imaging sessions advantageously minimize patient discomfort. Many patients are uncomfortable lying still during long imaging periods. While an extended imaging period at one imaging angle is not extremely burdensome, when images from many different imaging angles are required to generate a tomographic image, the duration of an entire imaging session can prove to be onerous. In these cases, adding even a few seconds to each imaging period to achieve threshold photon levels can increase patient discomfort appreciably.
Third, imaging systems are relatively expensive diagnostic tools and therefore the cost of such systems is usually only justifiable where a large number of patients can be examined each day.
Among other things, image quality is related to the type of processing performed by the camera processor. Generally, gamma camera processors can be divided into two different types, analog and digital. On one hand, analog processors are relatively fast but are only capable of moderate to poor spatial and energy resolution. On the other hand, digital processors are relatively slow but are capable of extremely accurate spatial and energy resolution. Thus, despite speed restrictions, to ensure quality images suitable for diagnostic purposes, most camera processors are of the digital type.
Image quality is also related to the number of photons within the marker energy range which are absorbed by the crystal. Thus, image quality can be increased by either increasing the photon generating intensity of the radiopharmaceutical in the organ of interest or by increasing imaging period durations. Because imaging speed is important, to the extent possible, image quality is increased by increasing photon intensity.
Unfortunately, image quality and imaging speed are limited to a certain extent by camera processor capability. The number of absorbed photons processed by a processor will be referred to as count rate. As with any processor, a camera processor is only capable of performing a maximum number of calculations per second MCPS and therefore can only accommodate a maximum count rate MCR. The MCR can be calculated according to the following equation: ##EQU1## where #s is the number intensity signals which must be processed for each absorbed photon and CPS is the number of calculations per intensity signal which must be performed to identify both photon energy level and impact point. For example, assuming a PMT array including sixty-three PMTs arranged in nine columns and seven rows, each PMT generating an intensity signal every time a photon is absorbed, #s is sixty-three. Also, assuming a processor having an MCPS of ten million and assuming CPS is 16, according to Equation 1, the MCR would be approximately 10,000 per second (i.e. MCR=10,000,000/(63)(16)).
If the number of absorbed photons per second exceeds the maximum count rate, the processor can experience "pile up" and "dead time" during which data related to some absorbed photons is effectively lost. As the name implies, pile up occurs when a processor cannot process all received signals and therefore must warehouse some of the signals in memory until processor time for processing the warehoused signals can be allotted. In some cases the memory can become full at which time additional received signals can be lost. Dead time results from the processor either failing to recognize essentially simultaneously absorbed photons and ignoring one of the photons or detecting essentially simultaneously absorbed photons but processing light associated with both photons as a single photon thereby causing quantitative errors and image artifacts.
There are several ways to increase maximum count rate MCR and thereby reduce the duration of imaging sessions without reducing image quality. The most obvious way to increase count rate is to provide a more powerful processor (i.e. increase MCPS in Equation 1). Unfortunately this solution is expensive to implement. In addition, in existing systems it would be relatively difficult to accommodate a hardware modification as processors are typically designed to perform specific required functions.
Another solution to increase the maximum count rate is reduce the duration of each PMT intensity signal by using a clipping circuit. U.S. Pat. No. 4,455,616 is exemplary of a signal clipping method. According to most signal clipping methods, after signal intensity corresponding to a single event exceeds a specific threshold energy level, the processor stops integrating light corresponding to the event and begins integrating light corresponding to subsequent events, thus freeing up some processor time. Unfortunately, while signal clipping methods reduce the time required to process intensity signals, these methods do not reduce the number of intensity signals which must be processed for each absorbed photon and thus the effect which these methods can have on count rate is limited. In addition, with signal clipping methods some imaging accuracy is sacrificed.
One other solution for increasing the maximum count rate is to reduce the number of calculations which have to be performed for each absorbed photon (i.e. reduce the denominator in Equation 1). To this end, one particularly advantageous solution to increase count rate has been to reduce the signal subset (i.e. #s in Equation 1) which is processed for each absorbed photon. In other words, assuming once again a PMT array including sixty-three PMTs arranged in nine columns and seven rows, instead of processing all sixty-three intensity signals which result from an absorbed photon, only a subset (e.g. five) of the sixty-three intensity signals are processed to identify photon intensity and impact location.
Generally, the industry has used three different methods for selecting a subset of intensity signals for processing. A first method is to identify intensity signals which have intensities above a threshold intensity level and then to process only the identified signals to determine photon energy level and impact location. For example, where the PMT array consists of sixty-three PMTs, when a photon is absorbed, all sixty-three intensity signals are converted to digital intensity signals. However, perhaps only the five strongest digital intensity signals are processed. This method reduces the number of intensity signals that must be processed, but does not reduce the effects of pile up on required energy and position calculations.
According to a second method, the processor divides the PMT array into first and second equal subsets, the first subset including roughly the left half of the array and the second subset including roughly the right half of the array. A third subset including the PMTs which form approximately the central one half of the PMT array is also earmarked by the processor. During operation, the processor combines digital intensity signals in the first subset and combines digital intensity signals in the second subset generating first and second combined intensity signals. If a combined intensity signal is above a threshold intensity level, a binary logic trigger signal (i.e. a "1") is generated for that subset. The two logic signals are combined to select the intensity signals corresponding to one of the three subsets to process. For example, assuming an array including nine columns of seven PMTs each, the first subset might include columns 1 through 5, the second subset might include columns 5 through 9 and the third subset might include columns 3 through 7. In this case, if the logic signal from the first set is true and the logic signal from the second set is false, signals from the first set are processed. If the logic signal from the first set is false and the logic signal from the second set is true, signals from the second set are processed. If both logic signals are true, signals from the third set are processed. Unfortunately, while this method does reduce the "zone" of PMTs to be processed. It only cuts the number of intensity signals to be processed in half.
A third method for selecting a subset of intensity signals to be processed can reduce the processing zone even further. U.S. Pat. Nos. 5,508,524 and 5,576,547 are exemplary of this method. According to this third method, digital intensity signals for each PMT are sent to an array of comparators. The largest intensity signal is determined by the comparators and used to determined which subset of intensity signals to process to determine photon intensity and impact location. While this method allows for smaller processing zones than the second method described above, this method requires comparisons between all intensity signals and therefore requires a complex comparator and routing circuit. Furthermore, since the comparison is made between processed intensity signals, an absorbed photon anywhere on the scintillation crystal will cause processing circuits for all of the PMTs to be busy.
In addition to each of the shortcomings identified above, methods which select a subset of intensity signals to process also reduce the accuracy of both energy and position calculations to a degree. In effect, a degree of accuracy is sacrificed for imaging speed. Unfortunately, while less accuracy might be required in some instances to increase imaging speed, it is not required in all instances and might therefore be sacrificed for nothing. For example, while photon absorption rate during one imaging period or at one point during an imaging period might be such that the processor can only process five intensity signals for each absorbed photon, during a subsequent imaging period or at a subsequent point in time during an imaging period, the absorption rate might be less such that all intensity signals corresponding to each absorbed photon could be processed. In these cases, methods which automatically select a small zone of PMTs for imaging needlessly sacrifice image accuracy.
Thus, it would be advantageous to have a method and/or apparatus for use with a gamma camera processor which could facilitate a processor count rate necessary to accommodate absorbed photons while at the same time maintaining the highest possible image quality consistent with the count rate.